1. Field of the Invention
The present invention is directed to magnetic resonance imaging (MRI or nuclear magnetic resonance (NMR)) imaging devices operated with a pulse sequence wherein a read-out sequence including at least two relatively-perpendicular gradients are generated per scan for location coding following a spin excitation, wherein the resulting signals are digitized and written in the K-space in a raw data matrix, wherein an image is acquired from the raw data matrix by conducting at least two-dimensional Fourier transformation, and wherein at least one gradient coil is connected with a capacitor to form a series resonant circuit connected to a gradient amplifier, the gradient amplifier being driven according to a predetermined time function.
2. Description of the Prior Art
A method for operating a nuclear magnetic resonance tomography apparatus known as the echo planar imaging (EPI) method is disclosed in European application 0 076 054, corresponding to U.S. Pat. No. 4,509,015. In summary, the echo planar imaging method includes the generation of an RF excitation pulse which is made slice-selective by simultaneously generating a magnetic field gradient in a first direction. A phase coding gradient is generated in a second direction, and a read-out gradient consisting of a gradient pulse sequence changing in polarity from pulse-to-pulse is generated in a third direction. The nuclear magnetic resonance signal acquired under the read-out gradient is phase demodulated, and is conducted through a bandpass filter. The output of the filter is digitized at a sampling rate and for each gradient pulse, is written into a row of a raw data matrix in the k-space. An image matrix is derived from the raw data matrix by two-dimensional Fourier transformation, and an image is produced from the image matrix.
Further details of the echo planar imaging method are discussed below in connection with FIGS. 1-9 to assist in the explanation of a problem associated with that known method to which the improvement disclosed herein is directed.
The basic components of a conventional nuclear magnetic resonance tomography apparatus are shown in FIG. 1. Coils 1-4 generate a static, fundamental magnetic field in which, if the apparatus is used for medical diagnostics, the body of a patient 5 to be examined is situated. Gradient coils are provided for generating independent orthogonal magnetic field components in the x, y and z directions, according to the coordinate system 6. For clarity, only gradient coils 7 and 8 are shown in FIG. 1, which generate the x-gradient in combination with a pair of identical gradient coils disposed on the opposite side of the patient 5. Sets of y-gradient coils (not shown) are disposed parallel to the body 5 above and below the body 5, and sets of z-gradient coils (not shown) are disposed at the head and feet of the body 5 extending transversely relative to the longitudinal axis of the body 5.
The apparatus also includes an RF coil 9 which excites selected nuclei in the body 5 so that nuclear magnetic resonance signals are generated, and also serves to acquire the resulting nuclear magnetic resonance signals.
The coils 1, 2, 3, 4, 7, 8 and 9 bounded by a dot-dash line 9 represent the actual examination instrument. The instrument is operated by an electrical arrangement which includes a fundamental field coils supply 11 for operating the coils 1-4 and a gradient fields coils supply 12 for operating the gradient coils 7 and 8 and the further gradient coils.
Via a switch 19, the RF coil 9 can be connected to an RF transmitter 15, in an excitation mode, or to an amplifier 14 in a signal reception mode. The amplifier 14 and the transmitter 15 are a part of an RF unit 16, which is connected to a process control computer 17. The computer 17 is also connected to the gradient fields coils supply 12. The computer 17 constructs an image from the nuclear magnetic resonance signals, which is portrayed on a display 18.
A number of pulse sequences are known for operating the RF unit 16 and the gradient coils. Methods have prevailed wherein the image generation is based on a two-dimensional or a three-dimensional Fourier transformation. One such method is the aforementioned echo planar imaging method.
A pulse sequence used in the echo planar imaging method is shown in FIGS. 2-6. A radio-frequency excitation pulse RF, shown in FIG. 2, is generated which excites nuclei in a slice of the examination subject which is selected by a slice-selection gradient SS in the z-direction, shown in FIG. 3, and generated simultaneously with the pulse RF. The direction of the gradient SS is subsequently inverted, the negative gradient portion of SS canceling the dephasing of the nuclear spins which was caused by the positive portion of the gradient SS.
After excitation, a phase coding gradient PC and a read-out gradient RO are generated. There are various possibilities for the respective curves of these two gradients. A phase coding gradient PC is shown in FIG. 4 which remains continuously activated during the read-out phase. An alternative phase coding gradient PC' is shown in FIG. 5 which consists of individual pulses ("blips") which are activated upon the occurrence of each polarity change of the read-out gradient RO. Each version of the phase coding gradient is preceded by a dephasing in gradient PCV in the negative y-direction. The read-out gradient RO is activated with a constantly changing polarity, as a result of which the nuclear spins are alternately dephased and rephased, so that a sequence of signals S arises. After a single excitation, so many signals are required that the entire Fourier k-space is scanned, i.e., the existing information is adequate for the reconstruction of a complete tomogram. For this purpose, an extremely rapid switching of the read-out gradient RO with high amplitude is required, which cannot be achieved with square-wave pulses which are usually employed in NMR imaging. A standard solution to this problem is the operation of the gradient coil which generates the gradient RO in a resonant circuit, so that the gradient RO has a sinusoidal shape.
The nuclear magnetic resonant signals S which arise are sampled in the time domain, are digitized, and the numerical values acquired in this manner are entered into a measurement matrix for each read-out pulse. The measurement matrix can be viewed as a measured data space, and in the exemplary two-dimensional embodiment as a measured data plane, in which the signal values are measured on an equidistant network of points. This measured data space is usually referred to in nuclear magnetic resonant tomography as the k-space.
Data identifying the spatial derivation of the signal contributions, which is needed for image generation, is coded in the phase factors, with the relationship between the locus space (i.e., the image) and the k-space being mathematically representable by a two-dimensional Fourier transformation. Each point in the k-space (in this case the k-plane) is therefore representable by the relationship: EQU .intg.(k.sub.x,k.sub.y)=.intg..intg..zeta.(x,y)e.sup.i (.sup.k x.sup.x+k y.sup.y) dxdy, wherein EQU k.sub.x (t)=.gamma..intg..sub.o.sup.t G.sub.x (t')dt',k.sub.y (t)=.gamma..intg..sub.o.sup.t G.sub.y (t')dt',
wherein .gamma. is the gyromagnetic ratio, and .zeta.(x,y) is the spin density distribution taking the relaxation times into consideration.
In FIGS. 8 and 9, the positions of the acquired measured values are schematically illustrated by points on a k-space trajectory in the k-space (k-plane). FIG. 8 shows the case for the continuous gradient PC of FIG. 4, and FIG. 9 shows the case for the gradient PC' shown in FIG. 5 in the form of a series of blips. For the Fourier transformation, the values must lie in an equidistant network of points, which is not the case in the examples shown in FIGS. 8 and 9. The acquired measured values therefore cannot be directly utilized, and an interpolation of the measured values onto an equidistant network of points must be undertaken.
Extremely high gradient amplitudes are needed for location coding of the NMR signals in the EPI method. These high gradient amplitudes must be activated and deactivated in short time intervals (&lt;1 ms), so that the necessary information can be acquired before the NMR signal decays. Due to the inductance and resistance of the gradient coils, these requirements cannot be satisfied in practice with a gradient coil connected directly to a gradient amplifier, because a terminal power of approximately 5 MW would be required.
The above problem can be resolved by connecting the gradient coil with a capacitor to form a parallel resonant circuit, as described in European application 0 227 411 and U.S. Pat. No. 4,628,264. To make parasitic transient effects as short as possible, the resonant capacitor is charged to the required voltage before the actual measuring sequence begins. After the field-generating gradient coil is connected into the circuit, the parallel resonant circuit immediately oscillates at the resonant frequency determined by the capacitor. There is thus substantially no transient effect. The capacitor, however, must be charged to an extremely high voltage (several kV), which requires that the gradient amplifier be capable of handling such voltages.
A series resonant circuit for generating gradient currents for the EPI method is generally described in the article "Whole Body NMR Spiral-Scan Echo Planar Imaging (SEPI) Using Resonant Gradient Coil," Kim et al., Society of Magnetic Resonance in Medicine, 7th Annual Meeting, Book of Abstracts, pg. 1013, however, the manner by which the series resonant circuit disclosed therein is caused to resonate is not discussed.
European application 0 389 666 discloses a gradient coil operated in a series resonant circuit in which the transients is controlled so that the current integral from the time of activation (t=t.sub.0) to a defined time (t=t.sub.1) just becomes zero. The capacitor of the series resonant circuit is thus charged to the necessary voltage without the gradient amplifier having to supply this voltage.
A disadvantage in these known circuits is the relatively long rise time caused by the inductance and resistance of the circuit, as well as by the output voltage of the gradient amplifier. Because the transient effect occurs with a relatively high operating frequency, the frequency-dependent resistance of the gradient coil also has a high value, which significantly lengthens the rise time.